Laser apparatus for treatment of a cataractous lens

ABSTRACT

An apparatus for aiding the removal of cataracts in which an optical fiber delivers sufficient optical energy of the correct wavelength, pulse duration to achieve controlled non-thermal and non-acoustic dissolution of hard cataract tissue.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present application is a continuation U.S. patent application Ser.No. 17/104,522 which is a continuation of U.S. patent application Ser.No. 15/512,129; which is a national stage application, filed under 35U.S.C. § 371, of International patent application no. PCT/CA2015/050915;which claims priority from U.S. provisional patent application No.62/052,109, filed Sep. 18, 2014; the entireties of which are all herebyincorporated by reference.

FIELD

The present disclosure relates to methods and apparatuses for deliveryof laser radiation for therapeutic purposes directed to and within acataractous lens.

BACKGROUND

A cataract is a clouding of the lens of the eyes which prevents clearvision. Although most cases of cataract are related to the agingprocess, occasionally children can be born with the condition, or acataract may develop after eye injuries, inflammation and some other eyediseases. Treatment for chronic deterioration of lens tissues is one ofthe most frequently performed surgeries.

In conventional cataract surgery, the eye surgeon typically uses ahand-held metal or diamond blade to create an incision in the area wherethe sclera meets the cornea. The next step for the cataract surgery istypically to remove the front portion of the capsule to allow access tothe cataract. Once the capsule is opened a tool can be inserted to breakapart and disrupt the cataract prior to removal. Tools for breakingapart the lens include mechanical tools such as scalpels or forceps totear the tissue apart, and more recently tools containing ultrasonictransducers have been used to emulsify tissue prior to aspiration. Evenmore recently, devices have been proposed that use laser radiation tobreak-down tissue through heating effects or acousto-optically generatedultrasonic energy for phacoemulsification (an example is described inU.S. Pat. No. 6,083,192). Most recently, techniques have been adopted inwhich radiation from very short pulsed lasers that are not absorbed wellin eye tissue are focused inside the volume of the cataractous lens toachieve photo-disruption of the tissue prior to aspiration.

However, conventional approaches may have one or more shortcomings.Using only mechanical tools, it is usually difficult and time consumingto carefully tear the lens tissue apart without creating uncontrolledstresses in the adjacent tissue, such as tearing of the capsule.

Ultrasound tools used for the phacoemulsification technique are usuallyable to effectively and quickly disintegrate hard lens tissue prior toaspiration. Ultrasonic energy however typically exerts negative effectson the tissues, including mechanical, thermal and non-thermal effects.Thermal effects are caused by the conversion of ultrasonic energy intothermal energy. This can result in heating or burning of the cornea. Theultrasound is essentially a high frequency mechanical perturbation ofthe tissue which disrupts the lens structure. This however isaccompanied by acoustic cavitation of the tissue and the resultant shockwaves which can propagate and further perturb tissue centimeters awayfrom the transducer. Furthermore, the ultrasonic formation of freeradicals during the cavitation process can damage delicate endothelialcells on the back surface of the cornea with oxidative stress.Ultrasonic energy propagates very well in aqueous tissue and the use oftoo much ultrasonic energy can lead to significant undesirablecomplications in parts of the eye beyond the lens, such as the corneaand retina.

Conventional devices which use laser radiation to generate theultrasonic energy typically suffer from the same limitations. Suchapproaches typically involve coupling pulsed laser light into the lenstissue using fiber optics for the purpose of ionizing, heating orshockwave generation by optical interaction with the tissue or some partof the tool tip. Examples are described in U.S. Pat. Nos. 4,744,360,6,623,477, 5,843,071, 5,919,186, and 6,083,192.

With the advent of picosecond and femtosecond pulsed lasers, scientistsfirst observed photodisruption, a different ablation mechanism in whichthe concentrated electromagnetic field of the short pulses destroysmatter by pulling it apart on a sub-atomic level. Reacting to the strongfields, the electrons in the material become energized beyond theionization limit (an example is described in U.S. Pat. No. 5,656,186).This mechanism is often referred to as “cold ablation” or “multi-photonionization” and has been proven to enable extremely precise machining ofmany materials. Regardless, the effects of this process on biology areonly recently being considered and there is concern for biologicaldamage due to free radicals caused by exposure of tissue to this kind ofionization radiation. Picosecond and femtosecond pulsed lasers have beenapplied to cataract surgery. Typically the surgeon creates a precisesurgical plan typically using a sophisticated 3-D image of the eye. Aspart of the preparatory steps for commencement of the surgery, thesefemtosecond laser systems are able to partially disrupt soft cataractouslenses by transmitting through the transparent portions of the eye andfocusing within selected portions of the lens to segment the cataractinto smaller pieces, with the goal of reducing or eliminating the use ofultrasound energy for lens disruption, and thereby reduce the risk ofburning and distorting the incision in the cornea. Using the fs laser inthis step may reduce the required phacoemulsification time, but fsradiation is not innocuous; and typically does not transmit consistentlywith unclear or scattering tissues in the beam path before the focusinside the lens. Furthermore, in most practical applications other thanvery soft cataracts, additional phacoemulsification is needed tobreak-up the remaining lens tissue. An example is described in U.S.Patent Application Publication No. 2009/0137993.

SUMMARY

In some examples, the present disclosure provides a laser-operatedapparatus and technique for disruption of cataractous-lens tissue priorto removal.

In various examples of the present disclosure, impulsive heat depositionis utilized to achieve micro-disruption of the lens tissue whilereducing or minimizing propagation of the energy to tissues other thancataractous-lens. This may be achieved by providing a tool which can beinserted within the volume of the cataract while providing suitableconditions for impulsive heat deposition upon contact with the distalend of the tool.

In some examples, the present disclosure provides an instrument whichembodies its own means of irrigation and aspiration of liquid at thesite of the fragmentation, without interfering with or diminishing theeffectiveness of the phacoablation.

In some examples, the present disclosure provides a surgical instrumentwhich enables external manipulation of the output end of an opticalfiber inside the eye, which may be directed only on nearby cataractouslens tissue to be fragmented. The particular laser that emits from thefiber tip is selected for its wavelength, intensity and pulse durationwhich may achieve conditions suitable for rapid micro-disruption throughimpulsive heat deposition.

In some examples, the present disclosure provides an apparatus fordisruption of cataracts in lens tissue. The apparatus includes: a sourceof pulsed laser radiation, the source being controllable to select apulsing rate of the pulsed laser radiation; an optical waveguideconfigured to transmit the pulsed laser radiation from the source, theoptical waveguide being coupleable to the source at a proximal end ofthe optical waveguide to receive the pulsed laser radiation from thesource; the pulsed laser radiation being controlled to exhibitconditions at a distal end of the optical waveguide such that the lightintensity which exits the optical waveguide is sufficient to producemicrodisruption of the lens tissue by impulsive heat deposition, theconditions including: a wavelength in the range of about 2700 nm toabout 3300 nm, the wavelength being selected to match an absorption peakof at least one component of the lens tissue; wherein the wavelengthcauses the pulsed laser radiation to produce laser pulses having anenergy sufficient to cause, when the laser pulses are absorbed in avolume of the material irradiated by the laser pulses, superheatedtemperatures above a vaporization point of the at least one component ofmaterial contained in the laser irradiated volume; a pulse duration timein the range of about 10 ps to about 1 ns, the pulse duration time beingselected such that each pulse duration time is shorter than a timerequired for thermal diffusion out of the laser irradiated volume andshorter than a time required for a thermally driven expansion of thelaser irradiated volume; wherein the combination of selected pulseduration time and selected pulse energy is low enough to result in apeak intensity of each laser pulse below a threshold forionization-driven ablation to occur in the irradiated material; andwherein the conditions are selected to result in conversion of amajority of the energy contained in each laser pulse to ablation of thematerial in the volume with any residual energy being insufficient tosubstantially damage material surrounding the volume irradiated by thepulsed laser.

BRIEF DESCRIPTION OF THE DRAWINGS

Reference will now be made, by way of example, to the accompanyingdrawings which show example embodiments of the present application, andin which:

FIG. 1 illustrates an example procedure for micro-disruption ofcataracts lens tissue, in accordance with an example of the presentdisclosure;

FIG. 2 shows the absorption spectrum of water from visible to farinfrared (IR);

FIG. 3 shows photographs of an example micro-disruption process, inaccordance with an example of the present disclosure;

FIG. 4 is a schematic diagram illustrating an example apparatus forcontrolled micro-disruption of cataract tissue, in accordance with anexample of the present disclosure;

FIG. 5 is a schematic diagram illustrating another example apparatus forcontrolled micro-disruption of cataract tissue, in accordance with anexample of the present disclosure;

FIG. 6 is a schematic diagram illustrating another example apparatus forcontrolled micro-disruption of cataract tissue, in accordance with anexample of the present disclosure;

FIG. 7 is a schematic diagram illustrating another example apparatus forcontrolled micro-disruption of cataract tissue including user feedbackor positional control, in accordance with an example of the presentdisclosure;

FIG. 8 shows various example geometries for the output face of the fiberin example apparatuses for micro-disruption of cataract tissue, inaccordance with examples of the present disclosure; and

FIG. 9 is a chart showing example experimental measurements of ablationthreshold.

Similar reference numerals may have been used in different figures todenote similar components.

DESCRIPTION OF EXAMPLE EMBODIMENTS

Various embodiments and aspects of the disclosure will be described withreference to details discussed below. The following description anddrawings are illustrative of the disclosure and are not to be construedas limiting the disclosure. Numerous specific details are described toprovide a thorough understanding of various embodiments of the presentdisclosure. However, in certain instances, well-known or conventionaldetails are not described in order to provide a concise discussion ofembodiments of the present disclosure. Although the present disclosuredescribes certain equations and/or theories to aid in understanding, thepresent disclosure is not necessarily bound to any of the describedequations and/or theories.

Nanosecond and longer pulsed mid-IR lasers have been used for ablationof ocular tissue such as cornea, however conventionally it had beenwidely considered best practice to avoid the use of pulse durationsshorter than a nanosecond to avoid the potential of ionization effects(see, for example, H. J. Hoffman, W. B. Telfair, “Minimizing thermaldamage in corneal ablation with short pulse mid-infrared lasers” J.Biomed. Opt. 4.4 (1999): 465). A mechanism for laser ablation, usingimpulsive heat deposition, was described in U.S. Pat. No. 8,029,501(which is hereby incorporated by reference in its entirety) in whichrapid-heating by excitation of vibrational modes inside of tissue causesvaporization of the exposed tissue. This has been shown in a number ofstudies to display unique laser material removal properties. However,applications of this cutting mechanism to cataract surgery have beenlimited due to the strong absorption in the eye tissue which limitsoperation to surface tissue.

In contrast to some of the previous solutions discussed above, invarious examples of the present disclosure, the laser ablation occursinside the body with the fiber tip surrounded by and in contact withtissue and fluid in the eye. There is no free surface for ablated tissueto expand into. Instead, the hard lens tissue is disrupted and the smallfragments are dispersed into the surrounding fluid in the eye.

In various examples, the present disclosure provides a cataract removalsystem that may avoid the energy propagation issues of thephacoemulsification process, photo-acoustic laser based systems.

In some examples, the present disclosure describes an apparatusincluding a laser probe which, on contact, and internal to the body, canefficiently drive rapid dissolution of lens tissue by optical excitationof selected vibrational modes inside of the tissue's molecules ontimescales faster than heat diffusion to the surroundings. In someexamples, the present disclosure describes an approach for efficientlydisrupting hard cataract tissue while avoiding the issues of energypropagation into other tissue's of the eye. In some examples, thepresent disclosure may provide one or more advantages over theconventional approach in respect efficient disruption of very hard lensmaterial, such as one or more of: less thermal or acoustic energyexposure to adjacent tissue, with or without an adjacent free surface;delivery through a fiber optic probe with sizes possibly down to thehundred micro size; and avoidance of tissue ionization and oxidativestress due to free-radical formation.

FIG. 1 shows an example illustration of micro-disruption of cataractslens tissue 1. In the example shown, micro-disruption of the cataractouslens tissue 1 occurs when laser pulses of a certain duration, wavelengthand pulse energy, are coupled into an optical waveguide 12 and exits (ata light exit) from the distal end 16 of the optical waveguide 12, wherethe distal end 16 has been inserted at some point 7, into the eye anddirected inside of the ocular lens 1. The light emitted from the distalend 16 is strongly absorbed by vibrational modes of the exposedmolecules of the lens cells 3 and/or intercellular regions 8, that arein contact with the light exit of the waveguide 12 or within a distanceclose to the optical absorption depth 40 of the laser light inside thetissue 1. The optical absorption depth 40 is a measurement of the extentto which the laser light is absorbed by tissue and/or fluid in thevicinity of the distal end 16. The optical absorption depth 40 may bedependent on the parameters of the laser light and/or the opticalproperties of the tissue and/or fluid surrounding the distal end 16. Thecells 3 and/or intercellular regions 8 that are exposed to the emittedlight together may be referred to as the irradiated volume 5. The resultis micro-disruption of the lens cells 3 and/or the cell structure 4 ofthe lens 1 faster than thermal diffusion or shockwave propagationoutside the irradiated volume 5. The excited molecules result ineffective dissolution 6 of the hard cataract lens 1 in such a way thatthe energy typically neither heats surrounding tissue, nor ionizes theexcited tissue, and typically prevents propagation of the energy todistant parts of the eye such as the cornea 2 and/or the lens capsule 9.Operation of the example waveguide 12 is further explained below.

Recently discovered molecular dynamic behavior of water molecules, insolution or bound to proteins and other molecules that comprise livingtissue, present a pathway to a laser-tissue interaction that isdifferent from prior mechanisms of mechanical, acoustic, or laserinduced breakdown, and that may provide advantages over conventionalapproaches. Example conditions suitable to produce this effect areprovided in the present disclosure. The selected combination, asdiscussed in greater detail below, of short pulse duration, wavelengthand pulse energy, pulse repetition frequency is delivered at the distalend of an optical waveguide.

The wavelength of the laser radiation should be strongly absorbed in thetissue, by transfer to vibrational modes. By targeting a strong peak inthe vibrational spectrum, such as the ubiquitous OH-stretch region ofH₂O, the vibrational modes may quickly absorb the electromagneticradiation and may effectively localize optical energy to micron scaledeep sections of the exposed tissue. This is illustrated by FIG. 2 ,which shows the absorption spectrum of water from visible to far-IR. Themaximum absorption occurs around 3000 nm where a broad peak correspondsto the OH-stretching vibrational modes of liquid water molecules betweenabout 2700 and 3300 nm. The spectrum also shows the resonance conditionsbetween the OH-stretch and other vibrational modes such as the OH bendand Intermolecular modes. Other absorption peaks, for example around theOH-bend at about 6000 nm, may also be used. In examples disclosedherein, the broad OH-stretch peak, in the range of about 2700 nm toabout 3300 nm, particularly around 3000 nm, are used since it may bemore effective and/or practical to produce laser light at thiswavelength range. Generally, in order for the ablation mechanismdescribed herein to be effective, the laser light should be selected tomatch a strong absorption wavelength of water or the tissue.

Subsequently, wavelengths in the mid infrared have an increasedthreshold for photo-ionization effects due to their lower photonenergies compared to near-IR, visible or UV lasers. Ionization oftissue, a mechanism that has its own intensity threshold forphoto-disruption, is an undesirable consequence which may be avoided byexamples of the present disclosure. The mechanism described hereintypically cannot be achieved at lower wavelengths, for example belowabout 1500 nm, where the multi-photon ionization occurs at thresholdslower than the requirements for micro-disruption through vibrationalexcitation of the material.

The pulse duration of the laser radiation should also be carefullyselected, as it dictates the minimum timescale at which energy isabsorbed and redistributed. Slow mechanisms of energy redistributionfrom optical excitation include thermal diffusion (many nanosecondtimescales) and shockwave emission (timescale>1 ns) that occur ontimescales orders of magnitude slower than fast mechanisms of energyredistribution such as avalanche ionization and vibrationalredistribution that occur on the femto-picosecond timescale (see, forexample Rafael R. Gattass & Eric Mazur, Femtosecond laser micromachiningin transparent materials. Nature Photonics 2, 219-225 (2008)). The rateof transfer of excited energy between vibrational modes in the presenceof water occurs on a particularly fast timescale compared to othermolecules (typically femtosecond to picosecond timescale) due to strongresonant coupling with lower frequency vibrational modes in the solvent.If the volume of excited tissue is large enough, e.g. micron scale, thetime required for diffusion of temperature or pressure gradients is muchlarger than the time required for those same gradients to disrupt thecellular structure of the tissue. In other words, this micro-disruptionis a process in which electromagnetic radiation drives theintra-molecular vibrations of the molecules in the tissue that quicklyand efficiently achieve molecular rearrangement (withoutphoto-ionization) and ultimately cellular scale mechanical motionsfaster than the energy can escape the irradiated volume as heat orshockwave.

A certain amount of pulse energy must be absorbed by a given volume oftissue to achieve the non-thermal and non-acoustic micro-disruptioneffect. Laser pulses in the picosecond time regime may be suited fordelivering the required energy to the tissue on this timescale whileavoiding peak intensities that would result in ionization. Ifinsufficient energy is delivered during the exposure of the laser pulse,the absorbed energy will dissipate as heat on thermal relaxationtimescales and the micro-disruption effect will not occur. If too muchenergy is delivered during the given pulse duration, the electromagneticfield intensities will begin to overcome the forces binding electrons totheir molecules and result in catastrophic photo-ionization of thetissue.

The micro-disruption threshold has been observed experimentally withpicosecond pulses and the effects of repeated exposure to belowthreshold optical excitation has been found to manifest themselves asmelting or burning of the tissue, whereas at above threshold opticalexcitation micro-disruption can be clearly observed. Above thethreshold, the tissue is disrupted with little, negligible orpractically no residual thermal effects.

Since the micro-disruption process may be less than 100% efficient thepulsing rate should also be considered. Individual laser pulses shouldhave sufficient energy to drive micro-disruption while allowing timebetween pulses for any residual energy left behind to dissipate beforethe next pulse of energy arrives, so as to reduce or preventaccumulation of the residual energy sufficient to drive other mechanismsof tissue damage such as increased temperature or shock waves. Laserrepetition rates in the 10-100 000 Hz range may enable average powerssuitable for fast tissue disruption with sufficient time between pulses.Bursts of multiple pulses at faster repetition rates may not satisfy thecriteria if sequential pulses that are below the energy threshold formicro-disruption are absorbed in the same volume at time intervalslonger than the relaxation time of the excited vibrational modes.

In the case of lens tissue, this photo-mechanism is enhanced by thecellular structure of the eye in which long, thin, transparent cells,with diameters typically between 4-7 microns and lengths of up to 12 mmare trapped in a regular pattern in shell like formations around thenucleus of the lens (as described in, for example, Biological glass:structural determinants of eye lens transparency, Phil. Trans. R. Soc.B. 2011 366 1250-1264). The majority of cells comprising the lens have aflattened hexagonal structure and are aligned into regular rows.Interdigitations are evident at the edges along the length as well as atthe ends of the fiber-like cells and act as an interlocking mechanism tomaintain the alignment of the cellular structure, which gives the lensits transparent optical properties in the visible spectrum. In the spaceseparating the cells, water and cell membrane proteins act to create afluid channel for cell hydration (as described in, for example,Gutierrez D B, Garland D, Schey K L. Spatial analysis of human lensaquaporin-0 post-translational modifications by MALDI mass spectrometrytissue profiling, Exp. Eye Res., 93:912-920, 2011). By selectivelyexciting the water molecules between cells and those on the surface ofthe proteins, it is possible to unravel the interlocking structure ofthe lens tissue so that the cells or portions of cells are easilydissolved into the fluid of the anterior portion of the eye.

FIG. 3 shows photographs of an eye while undergoing an examplemicro-disruption of tissue, in accordance with examples of the presentdisclosure. FIG. 3 a ) shows the cataract tissue of a human eye incontact with the distal tip of a 0.5 mm diameter solid sapphire fiberinto which pulses of 3000 nm, 400 ps, 500 uJ laser radiation energy arecoupled at a pulsing rate of 1 kHz. FIGS. 3 b ) and 3 c) show thevisible effect after exposure to several seconds of laser radiationdelivered to tissue that has come in contact with the distal tip of thefiber. The portions of the lens that were exposed through contact withthe distal tip of the fiber can be seen to scatter the light which isotherwise transmitted by neighboring tissue. FIG. 3 d ) shows the eyeafter complete disruption of the anterior portions of the lens as shownby the lack of reflected light.

In some examples, a picosecond pulsed (<1 ns) laser with wavelengthscorresponding to an absorption peak in the vibrational spectrum of water(around 3000 nm) and pulse energy E_(pulse), is coupled into a opticalwaveguide or fiber optic whose output aperture has an area of A anddirected inside the volume of a cataract such that the tissue which isdirectly in contact with the fiber tip can be exposed to lightintensities I=E_(pulse)/A which exceed the threshold required formicro-disruption of the targeted lens tissue. This intensity thresholdmay vary somewhat based on tissue characteristics, such as the tissuetype and in the case of cataracts, the age and/or hardness of thecataract. A lower limit for the intensity threshold may be approximately0.25 J/cm², as determined by experiment, example results of which areshown in FIG. 9 .

FIG. 9 shows example results of a measurement of the ablation thresholdusing 400 picosecond pulses from a 200 micron diameter fiber submersedin pure liquid water. The acoustic signal produced by the laser'sinteraction with the water is plotted versus the laser fluence. A changein behavior is seen at the ablation threshold near 0.25 J/cm². Acalculation of the pulse energy needed to vaporize the volume of waterexcited by the laser gives a similar result of 0.25 J/cm² for theablation threshold.

The upper limit for the intensity threshold may be determined by thephoto-ionization threshold, which is dependent on pulse duration. At thewavelength of about 3000 nm, the minimum pulse duration may be selectedto be about 10 ps to avoid ionization effects, and the maximum pulseduration may be selected to be about ins to avoid shock wave propagationin this tissue type. For a minimum pulse duration of about 10 ps, theupper limit for the intensity threshold may be experimentally determinedto be about 1 J/cm².

As an example, the fiber diameter, 2r, can be chosen to be about 0.5 mm.This fiber diameter was found in some cases to be a suitably large fiberdiameter for the selected pulse energy and intensity thresholds (asdiscussed above). In other examples where greater laser energy isselected, a larger fiber may be used. In this example, an intensityequal to the minimum ablation threshold is chosen thus requiring, forthis example, a pulse energy greater than

$E_{pulse} = {{I_{threshold} \times A} = {{0.25\frac{J}{{cm}^{2}} \times \times \left( {0.5/2} \right)^{2}} = {491 \times 10^{6}J}}}$

or approximately 0.5 mJ. In another example, the fiber diameter, 2r,could be chosen to be about 0.2 mm thus requiring a pulse energy at thelight exit at the distal tip of the fiber greater than

$E_{pulse} = {{I_{threshold} \times A} = {{0.25\frac{J}{{cm}^{2}} \times \times \left( {0.2/2} \right)^{2}} = {78 \times 10^{6}J}}}$

or approximately 0.08 mJ.

The equations presented above are illustrative and are not intended tobe limiting. The generalized form of this equation may be used todetermine the lower limit for the pulse energy required, for any givenfiber diameter. The upper energy limit may be found by experimentallydetermining the energy at which ionization damage occurs.

FIG. 4 illustrates an example of the present disclosure. A source oflaser pulses 10 is controlled by a signal 41, from a user input device11 (e.g., a computing device, a controller or a processing unit) and ameans for coupling the laser light into an optical waveguide or fiber12. A handle or fixture 13 is provided that allows insertion and controlof the distal tip 16 of the fiber 12 inside the lens portion 1 of thehuman eye for the purpose of cataract surgery. In some examples, aportion 13 a of the apparatus that comes in contact with the tissue maybe replaceable or re-useable. For example, the portion 13 a of theapparatus may be a single-use assembly, or a re-useable assembly thatmay be detached for sterilization and re-attached for repeated use. Theoptical output 14 from the distal tip 16 meets the conditions necessaryfor controlled micro-disruption of the exposed cataract tissue, forexample as discussed above.

FIG. 5 shows another example in which the optical fiber may be inserteddirectly or in combination with irrigation and or aspiration into thecataractous lens. In FIG. 5 , the source of laser pulses 10 iscontrolled by the user input device 11 in combination with a means ofirrigation 17 and a means of aspiration 18 (which are in turn alsocontrolled by the input device 11, via a control circuit 21, forexample). The source of laser pulses 10, the irrigation means 17 and theaspiration means 18 are coupled into a flexible, detachable, re-useableor disposable tool assembly 19 that allows insertion and control of thedistal tip 20 of the tool assembly inside an ocular lens 1 to achievecontrolled micro-disruption of the cataract tissue at the distal tip 20,as described above. One or more output channels for irrigation 51 andone or more input channels for aspiration 52 accompany the fiber opticdistal tip 16 to the disrupted lens material, which can be irrigatedand/or aspirated in a controlled manner with little or no loss or suddenchange of intraocular pressure.

In another example, the laser output may be controlled by a controllerexecuting an algorithm which receives inputs from one or more sensorsmonitoring variables such as the position and angle of the distal tip ofthe fiber, back scattered light emitting from the proximal end of thefiber, mechanical feedback (e.g., using a force sensor), acoustic and/orthermal conditions at or near the distal end of the fiber. The controlalgorithm may attempt to prevent accidental damage to surroundingtissues by shutting off the laser, when inputs from the one or moresensors indicate that surrounding tissues may be damaged. For example,one or more sensors may sense a temperature indicative of possibletissue damage (e.g., temperatures above a preset threshold). Othersensors, such as optical spectroscopic sensors or mass spectroscopicsensors may also be used to detect possible tissue damage. The one ormore sensors may also sense position of the distal tip 16 (e.g., usingaccelerometers or other suitable position sensors, such as 3D infraredtracking) to detect whether the distal tip is outside the expectedablation area. Generally, the sensor(s) may send appropriate signal(s)to the controller whenever the sensor(s) detects that conditions (e.g.,temperature, distal tip position, etc.) indicate a possible risk oftissue damage, and the controller may shut off the laser accordingly.The preset threshold(s) for the sensor(s) to indicate possible risk maybe preset to be lower than the threshold value(s) at which actual tissuedamage will occur, to factor in a safety margin.

In some examples, the control algorithm may be supplied with a 3D map ofthe boundaries of the lens (e.g., from prior imaging of the lens),enabling the controller to monitor the position of the tip and turn offthe laser outside of the predetermined lens boundaries so as to avoiddamage to surrounding tissues such as the capsule which should not bedisrupted or removed.

FIG. 6 shows an example including the use of sensors as described above.The source of laser pulses 10 is controlled by a signal 22 from acontrol circuit 21 (e.g., implemented in a controller, such as acomputing device) which receives inputs from one or more sensorsmonitoring variables such as a signal 53 (e.g., from a position and/ororientation sensor, such as an accelerometer) indicating the positionand angle of the distal tip 16 of the fiber, back scattered lightemitting from the proximal end of the fiber (e.g., detected by anoptical sensor 23 connected by a directional or wavelength dependantmeans 24 of coupling light into the optical fiber 12), a mechanicalfeedback signal (e.g., from a force sensor), and signals indicatingacoustic and/or thermal conditions at or near the distal end 16 of thefiber 12. When the control circuit 21 determines that the receivedsignals from one or more sensors indicate surrounding tissues 26 may beaccidentally damaged, the control circuit 21 shuts off the laser to thefiber 12. The control circuit 21 may also receive inputs, such asincluding a 3D map 25 of the boundaries of the lens (e.g., acquired inadvance by a suitable imaging technique) which allows the controlcircuit 21 to compare the position of the tip 16 (e.g., as indicated bythe position and/or orientation signal 23) with a preset boundarydefined in the 3D map 25, and prevent emission of the laser light atpositions outside of the predetermined boundaries so as to avoidaccidental damage to surrounding tissues 26.

In some examples, as shown in FIG. 7 , the control algorithm implementedby the control circuit 21 may also supply a control signal 54 to anactuator (e.g., a motor) of the handle or fixture of the fiber assemblyto control the position of the distal tip 16 and/or provide feedback tothe user in some way (e.g., tactile, audio or visual feedback). Similarto that described above, the location of the distal tip 16 of the fiber12 is monitored (e.g., using a position and/or orientation sensor thatprovides a position and/or orientation signal 53 to the control circuit21) and the position of the distal tip 16 is restricted to apredetermined volume which contains the lens material, so as to avoidaccidental exposure of surrounding tissues 26 (such as the capsule whichshould not be disrupted or removed) to laser radiation.

In some examples, the fiber is made of a relatively hard IR transmittingmaterial, such as sapphire (other suitable materials may includediamond, ZBLAN, YAG, etc.), and may have a tapered, curved or angled tipor any combination thereof, which may enhance the ease of use during thecataract disruption procedure.

FIG. 8 shows various example geometries for the output face at thedistal tip of the fiber. Some conditions which may be imposed by thesegeometries on pulse energy requirements to achieve the threshold ofselective micro-disruption are discussed below.

In the case of a cylindrical waveguide 61 with parallel walls and adiameter 28 of 2r the required pulse energy E_(pulse) needed to achievethe necessary conditions for micro disruption can be determined asfollows:

E_(pulse)≥I_(threshold)·πr², where I_(threshold) is the thresholdintensity of the micro-disruption process.

For a tapered waveguide 62, with an output aperture having diameter 29of 2r′ the required pulse energy would likewise be determined byE_(pulse)≥I_(threshold)·πr′² and the tapered angle 30, α, should be lessthan the critical angle for total internal reflection,

${\alpha < {\arcsin\left( \frac{n_{2}}{n_{1}} \right)}},$

where n₁ and n₂ are the index of refraction of the waveguide materialand the surroundings respectively.

For a waveguide with an angled output surface 63, of radius r, therequired pulse energy would likewise be determined byE_(pulse)≥I_(threshold)·πr²/sin θ, where the tip angle 32, θ, must begreater than the critical angle for total internal reflection

$\theta > {{\arcsin\left( \frac{n_{2}}{n_{1}} \right)}.}$

For a waveguide with a conical output surface 64, of radius r having anangle 34 of θ and cone length 35 of h, the required pulse energy wouldlikewise be determined by E_(pulse)≥I_(threshold)·πr(√{square root over(h²+r²)}), where the tip angle must be greater than the critical anglefor total internal reflection

$\theta > {{\arcsin\left( \frac{n_{2}}{n_{1}} \right)}.}$

A curvature of the distal portion of the fiber 65 can be useful, so longas the radius of curvature 36 does not exceed mechanical limits of thefiber itself, or cause loss of light propagation due to bending losses.

As used herein, the terms “comprises” and “comprising” are to beconstrued as being inclusive and open ended, and not exclusive.Specifically, when used in this specification including claims, theterms “comprises” and “comprising” and variations thereof mean thespecified features, steps or components are included. These terms arenot to be interpreted to exclude the presence of other features, stepsor components.

The embodiments of the present disclosure described above are intendedto be examples only. The present disclosure may be embodied in otherspecific forms. Alterations, modifications and variations to thedisclosure may be made without departing from the intended scope of thepresent disclosure. While the systems, devices and processes disclosedand shown herein may comprise a specific number of elements/components,the systems, devices and assemblies could be modified to includeadditional or fewer of such elements/components. For example, while anyof the elements/components disclosed may be referenced as beingsingular, the embodiments disclosed herein could be modified to includea plurality of such elements/components. Selected features from one ormore of the above-described embodiments may be combined to createalternative embodiments not explicitly described. All values andsub-ranges within disclosed ranges are also disclosed. The subjectmatter described herein intends to cover and embrace all suitablechanges in technology. All references mentioned are hereby incorporatedby reference in their entireties.

1. An apparatus for disruption of cataracts in lens tissue enclosed byat least one of aqueous humour or vitreous humour, the apparatuscomprising: a source of energy consisting of pulsed laser radiation, thesource being controllable to select a pulsing rate of the pulsed laserradiation; an optical waveguide configured to transmit the pulsed laserradiation from the source to the cataracts in the lens tissue, theoptical waveguide being coupleable to the source at a proximal end ofthe optical waveguide to receive the pulsed laser radiation from thesource; the pulsed laser radiation being controlled to exhibitconditions at a distal end of the optical waveguide such that the lightintensity which exits the optical waveguide is sufficient to producemicrodisruption of the lens tissue by impulsive heat deposition, theconditions including: a wavelength in the range of about 2700 nm toabout 3300 nm, the wavelength being selected to match an absorption peakof at least one component of the lens tissue; wherein the wavelengthcauses the pulsed laser radiation to produce laser pulses having anenergy sufficient to cause, when the laser pulses are absorbed in avolume of the material irradiated by the laser pulses, superheatedtemperatures above a vaporization point of the at least one component ofmaterial contained in the laser irradiated volume; and a pulse durationtime in the range of about 10 ps to about 10 ns, the pulse duration timebeing selected such that each pulse duration time is shorter than a timerequired for thermal diffusion out of the laser irradiated volume andshorter than a time required for a thermally driven expansion of thelaser irradiated volume to avoid shock wave propagation in the lenstissue; wherein the combination of selected pulse duration time andselected pulse energy is low enough to result in a peak intensity ofeach laser pulse below a threshold for ionization-driven ablation tooccur in the irradiated material; wherein the conditions are selected toresult in conversion of a majority of the energy contained in each laserpulse to ablation of the material in the volume with any residual energybeing insufficient to substantially damage material surrounding thevolume irradiated by the pulsed laser; and wherein, the apparatusfurther comprises a handle to insert and control the position of thedistal end of the optical waveguide inside the lens tissue.
 2. Theapparatus of claim 1 wherein the optical waveguide is an optical fibermade from an infrared-transmitting optical material.
 3. The apparatus ofclaim 2 wherein the optical fiber is made of a material selected from:sapphire, diamond, ZBLAN, or YAG.
 4. The apparatus of claim 2 whereinthe distal end of the optical fiber is curved, tapered, conical, orangled.
 5. The apparatus of claim 2, wherein the distal end of theoptical fiber is cylindrical.
 6. The apparatus of claim 1 furthercomprising one or more channels for irrigation and aspiration incombination and in close proximity to the distal end.
 7. The apparatusof claim 1 wherein the wavelength is selected to be about 2750 nm, andwherein the pulse duration time is less than about 5 ns.
 8. Theapparatus of claim 1, wherein the controller is configured to determinepossible unwanted exposure to laser radiation based on the trackedposition of the distal end of the optical waveguide relative to a threedimensional map indicating boundaries of the lens tissue, and thecontroller is further configured to reduce or cease laser radiation fromthe source in response to received signals indicating possible unwantedexposure to laser radiation.
 9. The apparatus of claim 1, wherein thecontroller further receives signals from an optical sensor detectingback scattered light that is transmitted from a vicinity of the distalend of the optical waveguide to a proximal end of the optical waveguide.10. The apparatus of claim 1, wherein the handle comprises a controller,and one or more sensors for tracking a position of the distal end of theoptical waveguide inside the lens tissue; and wherein the controller isconfigured to send signals to the handle to cause the handle to providea feedback based on the tracked position of the distal end of theoptical waveguide.